There are known whole body MRI magnets (super-conductive, resistive iron core magnets, and permanent magnets), which produce the background B.sub.o field used in MRI. The useable imaging volume in these magnets is in the region where the field homogeneity is a maximum. This volume is located in the air space centrally located between field sources. Thus, typically, MRI magnets are designed to provide a homogeneous magnetic field in an internal region within the magnet, e.g., in the air space of a large central bore of a solenoid or in the air gap between the magnetic pole plates of a C-type magnet. A patient or object to be imaged is positioned in the homogeneous field region located in such air space. In addition to the main or primary magnet that provides the background magnetic field B.sub.o, the MRI system typically has gradient and rf coils, which are, used respectively for spatial encoding and exciting/detecting the nuclei for imaging. These gradient field and rf coils are typically located external to the patient inside the geometry of the B.sub.o magnet surrounding the central air space.
Prior art electromagnets such as described by Watson et al and Muller et al. and other prior art iron core magnets typically have a structural design to provide a maximum magnetic field strength at a large central air space. In addition, those types of the prior art magnets, of the iron core electro- or permanent type, have a substantial edge fringe field effect, which makes it difficult to image beginning immediately at the magnet edge or even proximal to the edge of the magnet due to lack of sufficient field homogeneity.
In U.S. Pat. No. 5,049,848 a magnet configuration for MRI mammography is disclosed. The magnetic structure 50 has a rectangular shaped magnet with at least two parallel magnetic source 5,6 connected by a ferromagnetic core flux path defining an air gap for imaging. A remote shimming C-shaped magnetic source is preferably used for shimming to decrease the front edge fringe effect of the magnetic structure 50 to create a relatively homogeneous field in the air gap beginning at the front edge for effective imaging.
Solenoidal MRI magnets (superconductive, resistive) as well as iron core C and E shape electromagnets or permanent magnets are known for imaging of the whole body and its extremities. However, such whole body MRI magnets are not generally well-suited for treatment of the patient with other modalities or for minimally invasive surgical procedures guided by real time MRI because of the limited access of the surgeon to the patient. This limited access results from the field producing means surrounding the imaging volume. Electromagnets of the C or E type iron core configuration have been designed to offer a partially open access to the patient, however, the access is still very limited with typical air gaps of only 40 cm between the pole pieces of a C type magnet. U.S. Pat. No. 5,378,988 describes a MRI system, which can provide access for a surgeon or other medical personnel, using a plurality of C-shape solenoidal magnets oriented to form an imaging volume in a central region of the magnets.
Another type of magnet specifically designed for interventional surgical guidance is General Electric's Magnetic Resonance Therapy device, which consists of two superconducting coils in a Helmholtz coil type arrangement. The air gap for this magnet is 58 cm, which typically permits access by one surgeon. None of those prior art magnets or MRI systems is ideal with regard to simultaneously offering real time imaging and fully open access to the patient. Many surgical procedures require three or more surgeons together with an array of supporting equipment and, thus, a fully open magnet configuration for a MRI system for interventional procedures is desirable. In addition, such open magnet configuration is desirable for patients that have claustrophobia.
Applications other than MRI have used magnets that produce a useful field region outside the magnet geometry. U.S. Pat. No. 4,350,955 describes means for producing a cylindrically symmetric annular volume of a homogeneous magnetic field remote from the source of the field. Two equal field sources are arranged axially so that the axial components of the fields from the two sources are opposed, producing a region near and in the plane perpendicular to the axis and midway between the sources where the radial component of the field goes through a maximum. A region of relative homogeneity of the radial component of the background field B.sub.r may be found near the maximum. The large radial field is generally denoted as the B.sub.o background field in MRI applications. See also, J. Mag. Resonance 1980, 41:400-5; J. Mag. Resonance 1980, 41:406-10; J. Mag. Resonance 1980, 41:411-21. Thus, two coils producing magnetic dipole fields having opposing direction are positioned axially in a spaced relationship to produce a relatively homogeneous toroidal magnet field region in a plane between the magnets and perpendicular to the axis of cylindrical symmetry. This technology has been used to provide spectroscopic information for oil well logging but has not been used for imaging.
U.S. Pat. No. 5,572,132 describes a magnetic resonance imaging (MRI) probe having an external background magnetic field B.sub.o. The probe has a primary magnet having a longitudinal axis and an external surface extending in the axial direction and a rf coil surrounding and proximal to the surface. The magnet provides a longitudinal axially directed field component B.sub.z having an external region of substantial homogeneity proximal to the surface. Comparing this magnet geometry to that of U.S. Pat. No. 4,350,955, it has a background B.sub.o field with a cylindrically symmetrical region of homogeneity. However, this magnet described in the copending application provides such a field in the axial or z direction (i.e., longitudinal axis direction) whereas the other provides a background B.sub.o field in the radial or r direction (i.e., radial direction). Preferably, the B.sub.o field is provided by two magnets spaced axially and in axial alignment in the same orientation and wherein said region of homogeneity intersects a plane that is located between the magnets and that is perpendicular to the axis. For MR imaging, surrounding the primary magnet are r-, z- and .phi.- gradient coils to provide spatial encoding fields.
It is desirable to have new and better devices and techniques for biomedical MRI applications such as open magnet MRI systems for imaging while performing surgery or other treatments on patients or for imaging patients that have claustrophobia. It is also desirable to have portable devices and imaging techniques that could be applied to a wide variety of imaging uses.
Copending U.S. Ser. No. 08/695,174 filed on Aug. 8, 1996 describes a planar MRI system having an open magnet configuration comprising two pairs of planar pole pieces that produces a magnetic field having a substantial remote region of homogeneity.
Copending U.S. Ser. No. 08/869,009 filed Jun. 4, 1997 describes an open solenoidal magnet configuration comprises a pair of primary solenoidal coils and, located within the primary coil geometry, a bias coil system, the coils emitting an additive flux in the imaging region to generate a resulting field which provides a remote region of substantial field homogeneity.